1. Field of the Invention
The present invention generally concerns magnetic resonance tomography (MRT) as employed in medicine for the examination of patients. The present invention is in particular concerned with a method for determination of the B0 (basic magnetic) field, in particular given the use of fast MRT imaging methods such as, for example, TSE, EPI, SSEPI. A projection obtained of the B0 field (B0 field map) of the acquired slice enables the correction of image artifacts arising due to changes in the B0 field.
2. Description of the Prior Art
MRT is based on the physical phenomenon of nuclear magnetic resonance and has been successfully used as an imaging method in medicine and biophysics for over 15 years. In this examination method, the subject is exposed to a strong, constant magnetic field. The nuclear spins of the atoms of the subject, which were previously oriented randomly, thereby align. Radio-frequency energy can now excite these “ordered” nuclear spins to a specific oscillation (resonant frequency). This oscillation generates the actual measurement signal (RF response signal) in MRT, which is acquired by suitable receiver coils. Using non-homogeneous magnetic fields generated by gradient coils, the measurement subject can be spatially coded in all three spatial directions, which is generally designated as a “spatial coding”.
The acquisition of the data in MRT ensues in what is known as k-space (frequency domain). The MRT image in the image domain is linked with the MRT data in k-space by means of Fourier transformation. The spatial coding of the subject, which spans the k-space, ensues by means of orthogonal gradients in all three spatial directions. Differentiation is made between the slice selection (establishes an acquisition slice in the subject, typically the z-axis), the frequency coding (establishes a direction in the slice, typically the x-axis) and the phase coding (determines the second dimension within the slice, typically the y-axis).
A slice is thus initially selectively excited, for example in the z-direction. The coding of the spatial information in the slice ensues by a combined phase and frequency coding by means of these two already-mentioned orthogonal gradient fields which, in the example of a slice excited in the z-direction, are generated by the afore-mentioned gradient coils in the x-direction and y-direction.
A possible format to acquire the data in an MRT experiment is the method of echo-planar imaging (EPI). The basis of this method is to generate, after an individual (selective) RF excitation in a very short time span, a series of echoes in the readout gradient (Gx) that are associated by a suitable gradient switching (modulation of the phase coding gradient Gy) with various lines in the k-matrix. All lines of the k-matrix can be acquired in this manner with a single sequence pass (single-shot SS).
Such a single-shot method (SSEPI)—due to the speed of the acquisition of a slice—presently represents the dominant clinically applied method in MRT imaging with which functional MRT imaging (fMRT), perfusion imaging and diffusion imaging can be realized.
Variants of the echo planar technique differ only in how the phase coding gradients are switched, meaning how the data points of the k-matrix are sampled (scanned).
An EPI pulse sequence with a sinusoidally oscillating readout gradient and a constant phase coding gradient is shown in FIG. 5A. A constant phase coding gradient with sinusoidally oscillating readout gradient leads to a likewise sinusoidal sampling of the k-space, as is illustrated in FIG. 5b. The readout of the echo series must be concluded within a time span that corresponds (in terms of magnitude) to the decay of the transverse magnetization. Otherwise, the various lines of the k-matrix would be too significantly differently weighted, depending on their sequence of detection. In addition to this, interferences of local field inhomogeneities increase with increasing readout time. Due to the necessity of such high measurement speeds, the echo planar technique places very high requirements on the gradient system (in practice, for example, gradient amplitudes of approximately 25 mT/m are used; in particular to change polarities of the gradient field, significant energies must be converted in the shortest possible time; the switching times are, for example, in the range of ≦0.3 ms). Due to the large (in comparison with many other MR imaging techniques) length of the readout train, typically of 20-150 ms, the EPI method is sensitive to B0 field interferences. Temporally constant static effects influence the quality of the image data. The B0 field can, with good approximation, be assumed as constant over the length of the readout train for the following observations. In the application of the EPI technique, however, data typically are acquired over a longer time span of multiple minutes up to an hour and more. Over these time spans, fluctuations of the B0 field can occur due to external interferences (for example, elevators, street traffic, etc. in the proximity of the base field magnet of an MRT apparatus) as well as due to apparatus instabilities (a B0 field drift of the scanner is always known in principle). For example, given the long time series of data sets that are measured with EPI, fluctuations of the absolute value of the otherwise homogenous B0 field cause visible subject displacements in the phase coding direction (typical values: given a 128×128 pixel matrix, corresponding to approximately 10 Hz, one voxel—1.5 mm—displacement). In the case of functional imaging, such apparent subject movements are reliably removed from the measurement data as movement correction technique (Friston et al, Hum. Brain Map. 2: 165-189, 1995). Given the combination of EPI with contrast agent-supported methods (perfusion imaging), however, these effects can be corrected only with difficulty because the successive images are very different due to the contrast agent bolus flowing therethrough. This can lead to errors or given a low residual signal in the images, to outright impossibility of the image analysis. A similar problem occurs in diffusion imaging. The individual images are typically significantly different due to the strength of the diffusion coding as well as the coding direction. Furthermore, in particular given stronger diffusion coding, the information must be measured multiple times in order to actually obtain an acceptable ratio of signal to noise. The data of a single acquisition are unusable for a correction due to the high noise portion. If, for example, the B0 field now changes during the diffusion measurement, the individual acquisitions are displaced relative to one another, and the resulting image quality can be impaired.
Moreover, not only are temporal changes of the B0 field problematic, but also the variation of the absolute value of B0 across the sample or the patient to be examined. Normally, only B0 field deviations of the first and second order can be compensated by active and/or passive shimming. B0 field deviations of a higher order cause a residual curve of the B0 field of some 10 to 100 Hz in the homogeneity volume that cannot be corrected. Signals within the selected slice are possibly acquired off-resonance. In the case of echo planar imaging, this in turn leads to an apparent displacement of the measurement subject, meaning the planned or calculated image position (slice position) is typically incorrect by a few millimeters (up to centimeters). Problems with the interpretation and further use of the obtained images—for example given overlay with other measurement results for further planning of the examination and/or therapy—can result from this.
Conventionally in the case of echo planar imaging, artifacts due to temporal changes of the B0 field can be compensated under certain conditions. A prevalent method for this is known as “image matching”: a displacement or a drift of the subject in successive images is back-projected, by the rotation and/or translation being determined in which, for example, the difference of both images is minimal at the pixel level. In the case of imaging methods with significant contrast changes (such as, for example, diffusion imaging, perfusion imaging), however, this method frequently cannot be successfully used (given contrast changes, a comparison of the acquired data is increasingly more difficult). In addition, the image data of all individual acquisitions must exhibit a sufficiently high signal-to-noise ratio.
A further method for compensation of the effects occurring due to temporal changes of the B0 field is described in Durand (Durand et al., Journal of Magnetic Resonance in Medicine 46: 198-201, 2001): at each measurement (scan of the k-matrix), the phase of the middle (central) k-space line is averaged over the line, and this value is compared with the averaged phase of the preceding measurement (same slice) of the same line. Given B0 field fluctuations, averaged successive phases will be different. By comparison of the phases the B0 field, changes can be detected and the image can be correspondingly corrected and back-projected.
The method for correction proposed by Durand et al has also proven in practice to be only of limited capability. As already explained, with contrast agent-aided methods (perfusion imaging), the phase information is significantly altered by the (paramagnetic) contrast agent, such that a comparison of the averaged phases across a plurality of acquisitions collapses. Due to movements (for example, brain pulsation), diffusion imaging a priori leads to unknown displacements of the k-space data, such that a comparison of identical k-space lines is problematic. Moreover, in the diffusion coding, additional phase differences between the individual acquisitions of the comparison proposed by Durand et al of central k-space can be hampered or even can be completely impossible to devise. Finally, the received signal of individual acquisitions is increasingly severely noisy, in particular at higher diffusion coding strengths, such that no sufficiently precise phase information is available.